1. Technical Field
In general, the present disclosure relates to nuclear medical imaging. More particularly, the disclosure relates to Positron Emission Tomography (PET) imaging and accurate estimation of a timing coincidence window in a PET system.
2. General Background of the Invention
Nuclear medicine is a unique specialty wherein radiation emission is used to acquire images that show the function and physiology of organs, bones or tissues of the body. The technique of acquiring nuclear medicine images entails first introducing radiopharmaceuticals into the body—by either injection or ingestion. These radiopharmaceuticals are attracted to specific organs, bones, or tissues of interest. The radiopharmaceuticals produce gamma photon emissions, which emanate from the body and are then captured by a scintillation crystal. The interaction of the gamma photons with the scintillation crystal produces flashes of light or electromagnetic radiation in a different spectrum, which are referred to as “scintillation events.” Scintillation events are detected by an array of photo detectors (such as photomultiplier tubes (PMT) of avalanche photodiodes (APD)), and their spatial locations or positions are then calculated and stored. In this way, an image of the organ or tissue under study is created from detection of the distribution of the radioisotopes in the body.
One particular nuclear medicine imaging technique is known as positron emission tomography, or PET. PET is used to produce images for diagnosing the biochemistry or physiology of a specific organ, tumor or other metabolically active site. The measurement of tissue concentration using a positron emitting radionuclide is based on coincidence detection of the two gamma photons arising from a positron annihilation or coincidence event. When a positron is annihilated by an electron, two 511 keV gamma photons are simultaneously produced and travel in approximately opposite directions. Gamma photons produced by a coincidence event can be detected by a pair of oppositely disposed radiation detectors capable of producing a signal in response to the interaction of the gamma photons with a scintillation crystal. Coincidence events are typically identified by a time coincidence between the detection of the two 511 keV gamma photons in the two oppositely disposed detectors; i.e., the gamma photon emissions are detected virtually simultaneously by each detector. When two oppositely disposed gamma photons each strike an oppositely disposed detector to produce a time coincidence event, they also identify a line-of-response (LOR) along which the coincidence event has occurred. An example of a PET method and apparatus is described in U.S. Pat. No. 6,858,847, which patent is incorporated herein by reference in its entirety.
FIG. 1 is a graphic representation of a line of response. A coincidence event 140 occurring in imaged object mass 130 can emit two simultaneous gamma photons (not shown) traveling substantially 180° apart. The gamma photons can travel out of scanned mass 130 and can be detected by block detectors 110A and 110B, where the detection area of the block detector defines the minimum area or maximum resolution within which the position of an incident gamma photon can be determined. Since block detectors 110A and 110B are unable to determine precisely where the gamma photons were detected within this finite area, the LOR 120 connecting block detectors 110A and 110B can actually be a tube with its radius equal to the radius of block detectors 110A and 110B. Similar spatial resolution constraints are applicable to other types of detectors, such as photomultiplier tubes.
In commercial PET, detection of individual 511 keV gamma photons is accomplished by crystal photoelectric absorption whereby the gamma photon's energy is converted to light, or some other electromagnetic radiation having a different frequency, by scintillation crystals contained within a detector. The production of the electromagnetic radiation by the scintillation crystal is known in the art as a scintillation event. Typically, PET photon detectors (photo detectors) are shared amongst an area of crystals to allow a small number of photo detectors to support a larger number of scintillation crystals as shown in FIG. 2.
The detector scintillation crystals are not optimally coupled to the photo detectors in a 1:1 configuration since such an implementation is cost prohibitive for a commercial whole body system. Multiple scintillation crystals can be associated with one photo detector. For example, FIG. 2 illustrates a physical layout of photo detectors 0-8. The photo detectors 0-8 used in a PET block detector can be, for example, PMTs or APDs. In the example illustrated in FIG. 2, APDs are used; however, PMTs or other photo detectors can also be used. Each photo detector is numbered consecutively starting at APD0 at the top left corner of the block. Nine APDs are arranged for detecting light emitted by 100 scintillation crystals.
Ideally, each event detected can be correlated with a corresponding coincidence event; however, in real-world applications, coincidence events are often detected because of scattered gamma photons or other random events (randoms). Randoms result in spurious detection of coincidence events and thereby create noise in a finally rendered image. Therefore, the ability to differentiate randoms from true coincidence events is a long-felt and unresolved problem in the art.